Blood flow monitor with venous and arterial sensors

ABSTRACT

A technique is disclosed for determining blood flow in a living body by changing the thermal energy level in the venous blood flow path and determining temperatures in both the venous and arterial blood flow paths. Blood flow is calculated as a function of the change in energy level and the temperature differences in the venous and arterial blood flow paths.

PRIORITY CLAIM

[0001] This application claims priority from commonly owned copendingU.S. Provisional Application Serial No. 60/458,100, filed Mar. 26, 2003,and commonly owned copending Nonprovisional U.S. application Ser. No.10/364,773, filed Feb. 11, 2003.

INTRODUCTION

[0002] This invention relates generally to techniques for measuringblood flow in a body and, more particularly, to the use preferably ofone or more temperature sensors for measuring thermal energy changes inthe blood flowing through the heart and to the use of unique dataprocessing techniques in response thereto for determining cardiacoutput.

BACKGROUND OF THE INVENTION

[0003] While the invention can be used generally to measure blood flowat various locations in a body, it is particularly useful in measuringblood flow in the heart so as to permit the measurement of cardiacoutput. Many techniques for measuring cardiac output have been suggestedin the art. Exemplary thermodilution techniques described in thetechnical and patent literature include: “A Continuous Cardiac OutputComputer Based On Thermodilution Principles”, Normann et al., Annals ofBiomedical Engineering, Vol. 17, 1989; “Thermodilution Cardiac OutputDetermination With A single Flow-Directed Catheter”, Forrester, et al.,American Heart Journal, Vol. 83, No. 3, 1972; “Understanding Techniquesfor Measuring Cardiac Output”, Taylor, et al., BiomedicalInstrumentation & Technology, May/June 1990; U.S. Pat. No. 4,507,974 ofM. L. Yelderman, issued Apr. 2, 1985; U.S. Pat. No. 4,785,823, of Eggerset al., issued on Nov. 22, 1988; and U.S. Pat. No. 5,000,190, of John H.Petre, issued on Mar. 19, 1991.

[0004] A principal limitation in the quantification of cardiac output isthe existence of thermal fluctuations inherent in the bloodstream.Previous methods work with those fluctuations while observing theeffects of an input signal to calculate cardiac output. The inventiondescribed herein uses a differential measurement technique tosubstantially eliminate the effect of the thermal fluctuations,permitting the use of a minimal thermal input signal, which allowsfrequent or continuous measurements.

[0005] It is desirable to obtain accurate cardiac output measurements inan effectively continuous manner, i.e., several times a minute, so thata diagnosis can be achieved more rapidly and so that rapid changes in apatient's condition can be monitored on a more continuous basis than ispossible using current techniques. Moreover, it is desirable to obtaininstantaneous measurements of the cardiac output on a beat-to-beat basisto evaluate the relative changes which occur from beat to beat, as wellas to determine the presence of regurgitation.

BRIEF SUMMARY OF THE INVETION

[0006] In accordance with general principal of the invention, blood flowand/or cardiac output is determined rapidly, using a technique by whichan indicator substance, or agent, is introduced into the bloodstreambetween a pair of detectors. The detectors are sensitive to a parameterfunctionally related to the concentration or magnitude in thebloodstream of the selected indicator agent. The detectors arepositioned apart by a distance functionally sufficient to allow ameasurement to be made of the differential value of the selectedparameter as it exists from time-to-time between the two detectors. Theindicator agent, for example, may be a substance to change the pH of theblood, a fluid bolus carrying thermal energy, or a substance to change aselected characteristic of the blood, or the direct introduction ofthermal energy, or the like.

[0007] A determination is made of the difference in the values of theselected blood parameter as it exists at the two detectors, prior to theintroduction of indicator agent (i.e., the first differential value).The selected indicator agent is then introduced in a predeterminedmagnitude. Then again a determination is made of the difference invalues of the selected parameter as it exists at the locations of thetwo detectors (i.e., the second differential value).

[0008] Blood flow or cardiac output, depending on the specific locationof the detectors, can then be determined as a function of the differencebetween the first differential value and the second differential value.Because the ultimate measurement of blood flow or cardiac output isbased on the difference of the differences, the system operateseffectively with the introduction of the indicator agent in a very lowmagnitude. In turn, this allows measurements to be made rapidly so thateffectively continuous measurements are obtained.

[0009] In accordance with a preferred embodiment of the invention, forexample, cardiac output can be determined rapidly and with low levels ofthermal energy input. To achieve such operation, in a preferredembodiment, the technique of the invention uses a pair of temperaturesensors positioned at two selected locations within a catheter which hasbeen inserted into the path of the blood flowing through the heart of aliving body. The, sensors detect the temperature difference between thetwo locations. Depending on the location of the temperature sensors inthe circulatory system, the measured temperature difference varies overtime. It has been observed that when the temperature sensors are placedwithin the heart, e.g., so that one sensor lies in the vena cava, forexample, and the second in the right ventricle or pulmonary artery, thetemperature difference varies in a synchronous manner with therespiratory cycle.

[0010] Thus, in the preferred embodiment of the invention thetemperature difference over at least one respiratory cycle is measuredand averaged to provide an average temperature difference. Theaveraging, or integrating, action effectively eliminates, as aconfounding factor in the determination of cardiac output, the effect ofinstantaneous blood temperature fluctuations, such as cyclical,respiratory-induced fluctuations.

[0011] To make such determinations, an average temperature difference isfirst calculated over a time period of at least one respiratory cycle inwhich no thermal energy is introduced into the blood flow path. Thermalenergy of a predetermined and relatively low magnitude is thenintroduced into the blood flow path to produce a heating action thereinat a location between the two temperature sensors. Once the temperaturerise induced by the heating stabilizes, the average temperaturedifference between the two locations is again calculated fromtemperature difference measurements over a time period of at least onerespiratory cycle at the higher temperature level. The differencebetween the average temperature differences which occurs when thethermal energy is turned on, referred to as the rising temperaturechange, is determined. The difference between the average temperaturedifferences which occurs when the thermal energy is turned off, referredto as the falling temperature change, is similarly determined. Thecardiac output is calculated as a function of the thermal energy inputand the rising and falling temperature changes. Because a relatively lowlevel of thermal energy is used in making measurements, the overallsequence of determinations can be safely repeated multiple times perminute, for example, so that an effectively continuous, orquasi-continuous, determination of cardiac output is obtained.

[0012] Further, a pair of sensors may be introduced into the circulatorysystem at other locations in a subject. For example, a catheter may beintroduced into the venous side of the circulatory system to locate anenergy source (e.g.: a heater) in or near the vena cava or within theheart. A sensor may be located in the venous side of the circulatorysystem, upstream of the energy source, anywhere that temperature is notmaterially affected by output from the energy source. This sensor is onemeans for providing a reference signal that compensates for fluctuationsin bloodstream temperature introduced by factors other than the energysource. That is, the reference signal compensates for backgroundbloodstream temperature. The other sensor is introduced into thearterial side of the circulatory system anywhere it senses bloodtemperature as affected by the energy source. It may be introduced by asecond catheter, for example an arterial catheter. Blood flow and/orcardiac output may then be calculated as described above.

[0013] In accordance with a further embodiment of the invention, atemperature sensor that also acts as a source of thermal energy, e.g., athermistor, is positioned at a third location in the cardiac blood flowpath. Power is supplied to the sensor sufficient to elevate thetemperature of the sensor from a first temperature level to a secondtemperature level. In one embodiment of the invention, the temperatureof the sensor is changed from the first to the second level and ismaintained constant at said second level by varying the power that issupplied thereto. Such varying power is proportional to theinstantaneous flow velocity and, hence, assuming a constant flow area,is proportional to the instantaneous cardiac output. Measurement of thesensor heating power and the temperature increment at the sensor canthus be used to continuously effect a determination of the instantaneouscardiac output. Further, for example, when the sensor is placeddownstream at the outlet of one of the heart chambers, the variation inflow output over the cardiac cycle can be analyzed to provide anindication of the regurgitation characteristics of the heart outletvalve over the cardiac cycle. Moreover, such instantaneous cardiacoutput determination can be further refined to compensate forfluctuations in the temperature of the blood flowing through the heartby measuring the instantaneous temperature of the blood with anothertemperature sensor at a nearby location and appropriately taking intoaccount such temperature variations when determining the cardiac output.

[0014] In another application, both the continuous cardiac outputdeterminations and the instantaneous cardiac output determinations, asdescribed above, can be combined. Thus, three temperature sensors and asource of thermal energy can all be used in combination tosimultaneously provide an accurate and effectively continuousdetermination of time-averaged cardiac output, and a determination ofinstantaneous cardiac output at each instant of the cardiac cycle instill another application, two temperature sensors and a source ofthermal energy can be used in an appropriate sequence to provide theaveraged cardiac output determination and the instantaneous cardiacoutput determination.

DESCRIPTION OF THE INVENTION

[0015] The invention is described in detail with the help of theaccompanying drawings wherein:

[0016]FIG. 1 shows a simplified diagrammatic view of a human heart;

[0017]FIG. 2 shows a simplified diagrammatic view of a catheter usefulin the invention;

[0018]FIG. 3 shows a flow chart depicting steps in a process used in theinvention;

[0019]FIG. 3A shows a smooth temperature difference curve obtained inthe process depicted in FIG. 3;

[0020]FIG. 4 shows a flow chart depicting further steps in a processused in the invention;

[0021]FIG. 5 shows a flow chart depicting still further steps in aprocess of the invention;

[0022]FIG. 6 shows a graph depicting a temperature/time relation used inthe invention;

[0023]FIG. 7A shows a simplified diagrammatic view of another catheterused in the invention;

[0024]FIG. 7B shows a flow chart depicting steps in still anotherprocess of the invention;

[0025]FIGS. 8A, 8B and 8C show flow charts depicting modes of operatingthe process of the invention;

[0026]FIGS. 9A, 9B, and 9C show graphs of parameter relationships usedin the invention; and

[0027]FIG. 10 shows a graph useful for calibrating the flow values for acatheter used in the invention.

[0028]FIG. 11 is a diagram of another catheter configuration.

[0029] As can be seen in FIG. 1, which represents a human heart 10 in amuch simplified diagrammatic form, a flexible catheter 11 is insertedthrough the veins into the right atrium, or auricle, 12 of the heartand, thence, through the right ventricle 13 until the end of thecatheter resides in or near the exit, or pulmonary, artery 14 whichleads to the lungs. As is well known, blood flows (as represented by thearrows) from the input vein 15, i.e., the vena cava, into the rightatrium and right ventricle and thence outwardly to the lungs andsubsequently returns from the lungs into the left atrium 16, through theleft ventricle 17 and thence outwardly into the aorta 18.

[0030] In accordance with the embodiment of the invention, shown withreference to FIG. 1, temperature sensors, e.g., thermistors, are carriedby the catheter so that, when inserted as shown in FIG. 1, a firstsensor 19 is positioned at a location within the vena cava 15 or rightatrium 12 and a second sensor 20 is positioned at a location in or nearthe pulmonary artery 14.

[0031] For simplicity, the flexible catheter 11 is depicted in FIG. 2 inan extended condition with temperature sensors 19 and 20 at twodifferent locations for measuring temperatures T.sub.1 and T.sub.2,respectively. A power source 21 of thermal energy which is borne, orcarried, by the catheter 11 is positioned in the right atrium at alocation between sensors 19 and 20. In a particular embodiment, thecatheter-borne source is, for example, a coil of resistive wire placedon or embedded in the surface of catheter 11, to which an AC or a DCvoltage (not shown) at a controllable level is supplied so as togenerate thermal energy, i.e. heat. The magnitude of the thermal energycan be suitably controlled to insert a predetermined amount of thermalenergy at a selected time, which thermal energy is transferred to theblood flowing through the heart so as to raise its temperature. Theenergy source is positioned at a sufficient distance from the sensor 19that the latter is effectively thermally isolated from the site of thethermal energy source.

[0032] While the locations of the sensors 19 and 20 and the energysource 21 can be as shown in FIG. 1, alternative locations can also beused. Thus, the sensor 19 can be positioned in the vena cava 15, whilethe energy source 21 is located in the right atrium 12 and the sensor 20in either the right atrium or the right ventricle. Moreover, if sensor19 is positioned in the vena cava 15, the entire energy source 21, whichis normally elongated, need not be located in the right atrium and canhave a portion thereof in the vena cava and a portion thereof in theright atrium. Such source should preferably be at least partiallylocated in the right atrium. Further, sensor 20 may be positioned in theright ventricle near the pulmonary artery 14 or may be located in thepulmonary artery itself at or near the right ventricle.

[0033] The temperatures T₁ and T₂ at locations 19 and 20 upstream anddownstream, respectively, from the thermal energy source 21 aremonitored and processed appropriately by a digital microprocessor. Inaccordance with the invention, the instantaneous temperatures areobtained as the outputs T₁(t) and T₂(t) of the temperature sensors 19and 20, respectively. The outputs are connected to a differentialamplifier to generate an analog signal which is proportional to thetemperature difference ΔT(t)=T₁(t)−T₂(t) between them. The temperaturedifference signal ΔT(t) is digitized and sampled at selected timeintervals by an analog-to-digital/sampling circuit. The digitizedsampled temperature difference values and the known thermal energyvalues are supplied to a digital microprocessor which then suitablyprocesses the data to provide the desired cardiac output information.The processing stages used in the host microprocessor are implemented bysuitable programming of the microprocessor and are discussed below withthe help of FIGS. 3-6.

[0034] The source 21 of thermal energy is alternately turned on and off.If it is assumed that thermal stability is reached after each change andthat there is a substantially constant rate of blood flow, a stabletemperature difference can be measured in each case. The quantity ofblood flowing past the thermal energy source, i.e., the cardiac output,can be derived from such temperature difference measurements. However,such derivation is complicated by two factors which may affect themeasurement of blood flow. First, the rate of blood flow through theheart is not substantially constant but surges with each heartcontraction. Second, the temperature of the blood flowing through theheart is not constant but varies with each respiratory (breathing)cycle. In a preferred embodiment, the processing of the data takes suchfactors into account, as discussed below.

[0035] The process for determining cardiac output is performed in amicroprocessor 2 which in a first embodiment is programmed to respond tothe temperatures sensed at T.sub.1 and T.sub.2 and to perform the stepsdepicted in accordance with the flow charts shown in FIGS. 3-5. From aknowledge of such flow charts, it would be well within the skill ofthose in the art to appropriately program any suitable and known digitalmicroprocessor, such as a personal computer, to perform the steps shown.

[0036]FIG. 3 depicts a basic process, identified as Process I, which isused in the overall processing of temperature data for determiningcardiac output, as subsequently depicted in FIGS. 4 and 5. In the basicprocess shown in FIG. 3, a temperature difference as a function of timeΔT(t) is determined by a differential amplifier which responds to T₁(t)and T₂(t). Such differences may be effectively smoothed, or filtered, toproduce a smooth temperature difference curve, as shown in FIG. 3A,which varies as a function of time in a cyclic manner which dependsprincipally on the respiratory cycle of the person whose cardiac outputis being determined.

[0037] The periods τ₁, τ₂ . . . τ_(n) for each respiratory cycle aredetermined over n cycles. A characteristic of the temperature differenceat each cycle is determined. For example, such characteristic preferablyis the averaged temperature difference during each cycle (ΔTτ₁, ΔTτ₂ . .. ΔTτ₂ . . . ΔTτ_(n)).

[0038] (Alternatively, for example, the peak temperature differences maybe the determined characteristic.) These averaged temperaturedifferences (ΔT_(τn)) are added for the n cycles involved and aredivided by n to determine an averaged temperature difference per cycle(ΔT_(τn)). The use of Process I is depicted in the process steps shownin FIG. 4, identified as Process II.

[0039] As seen therein, the steps of Process I are first performed whenthe source 21 of thermal energy (i.e., a heater) is turned off and theaverage AT_(off) value per cycle is determined and suitably stored. Thesampling time at which such determination is made is depicted in FIG. 6as the sample time period S1.

[0040] The heater 21 is then turned on for a specific time period tosupply a known amount of power P to the blood flowing through the heartand, accordingly, the temperature of the blood flowing past the heaterrises and the temperature difference ΔT(t) rises over a transition, ordelay, rise time period, t_(R1), shown in FIG. 6 and designated as D1,after which the temperature difference generally stabilizes over asecond sample time period S2. As seen in FIG. 4, after the heater 21 isturned on and the temperature has stabilized, Process I is performed,again over n cycles, e.g., over the time period S2, and the averagedtemperature difference ΔT_(on) is determined with the heater turned onand is suitably stored. The heater is then turned off and thetemperature falls over a transition, or delay, fall time periodt.sub.F1, shown in FIG. 6 and designated as D2, generally to its formervalue.

[0041] Cardiac output is calculated using the averaged temperaturedifferences when the energy is off and the averaged temperaturedifferences when the energy is on, by the relationship:$F = \frac{P}{C_{p}\left( {{\overset{\_}{\Delta \quad T}}_{on} - {\overset{\_}{\Delta \quad T}}_{off}} \right)}$

[0042] where:

[0043] F=Flow

[0044] P=Power

[0045] C_(p)=heat capacitance

[0046]ΔT _(on)=average temperature for power on

[0047]ΔT _(off)=average temperature for power off.

[0048] As seen in FIG. 5, the steps of Process II are repeatedindefinitely for N data collection cycles, a data collection cycle beingdesignated as including the time periods S1, D1, S2, and D2, as shown inFIG. 6. For each data collection cycle the rise time temperaturedifference ΔTR between the averaged temperature difference ΔT_(on) at S2and the averaged temperature difference ΔT_(off) at S1 and the fall timetemperature difference ΔT_(F) between the averaged temperaturedifference ΔT off at S1 and the averaged temperature difference{overscore (ΔT)}_(on) at S2 are determined.

[0049] The flow, F_(R), is calculated for each data collection cyclefrom the known amount of power P introduced into the blood flow streamby the energy source, or heater 21, from the known heat capacitance ofblood, C_(P), and from the difference in the averaged temperaturedifferences {overscore (ΔT)}_(on) and ΔT _(off), which occurs over thedata collection cycle S1+D1+S2 in accordance with the followingrelationship:$F_{R} = \frac{P}{C_{p}\left( {{\overset{\_}{\Delta \quad T}}_{on} - {\overset{\_}{\Delta \quad T}}_{off}} \right)}$

[0050] In a similar manner, the flow F_(F) is calculated from P, C_(P)and the difference in the averaged temperature differences ΔT_(on) andΔT_(off) which occurs over the later portion of the data collectioncycle S1+D1+S2 in accordance with the following relationship:$F_{F} = \frac{P}{C_{p}\left( {{\overset{\_}{\Delta \quad T}}_{on} - {\overset{\_}{\Delta \quad T}}_{off}} \right)}$

[0051] F_(R) and F_(F) can be averaged to obtain the averaged flow({overscore (F)}) over one data collection cycle as shown in FIG. 6.$\overset{\_}{F} = \frac{F_{R} + F_{F}}{2}$

[0052] A suitable calibration constant can be used to adjust the valuesof F_(R), F_(F) and {overscore (F)}.

[0053] Accordingly, by using two temperature sensors 19 and 20, cardiacoutput can be determined several times a minute in accordance with FIGS.3-6, yielding an effectively continuous cardiac output value. Becausesuch measurements can be made using relatively low power levels, thedanger that the heart may be damaged by the introduction of thermalenergy is substantially eliminated.

[0054] It will be apparent that the foregoing technique, which has beendescribed in connection with the direct introduction of heat as anindicator agent and the measurement of temperature, can readily beperformed by those skilled in the art by using indicator agents whichaffect the pH of the blood or change other blood parameters.

[0055] The system can be located in the body of a subject as illustratedin the diagrammatic representation of the circulatory system shown inFIG. 11. The circulatory system illustrated there includes systemicvenous system 101, vena cava 150, heart 100, systemic arterial system102, pulmonary venous system 106 and the body's capillary system 216.The representation of the capillary system 216 is intended to indicatethat system anywhere in the body. The representation of the heart 100includes the right atrium (RA), the right ventricle (RV), the leftatrium (LA) and the left ventricle (LV).

[0056] An energy source 210 and a sensor 190 are located in the venousside 101 of the circulatory system of the subject by any suitableplacement method or means. For example, a catheter 110 may be introducedinto the vena cava 150 to locate the energy source 210 in or near thevena cava or within the heart 100 and the sensor 190 somewhere in thevenous system upstream of the energy source 210.

[0057]FIG. 11 shows the energy source 210 in the vena cava 150 whileFIG. 1 shows an energy source 21 in the right atrium 12. The energysources 21 and 210 are similar in function and may be similar inconstruction. Sensor 190 is located upstream of the energy source 210where it is not materially affected by output from the energy source,where it senses blood temperature unaffected by the energy source.Sensor 190 or other suitable reference provides a reference signal thatcompensates for temperature level and fluctuations in bloodstreamintroduced by factors other than the energy source 210. That is, thereference signal corresponds to background bloodstream temperature,unperturbed by the energy source 210. FIG. 11 shows one preferredlocation for sensor 190 in or near the vena cava and an alternatelocation more remote from the energy source. In another embodiment ofthe system a fixed resistor or other reference value that approximatesthe unperturbed blood temperature could be used, in lieu of the valueprovided by the sensor 190, to compensate for background bloodstreamtemperature. (This likely would reduce the accuracy of the compensationand its use would depend on the requirements of the user.) Sensor 190can be similar in construction and is similar in function to sensor 19shown in FIG. 1.

[0058] The sensor 200 is located in the arterial system 102 and may beintroduced by any suitable method or means. For example, an arterialcatheter may be used. The sensor 200 may be located anywhere on thearterial side 102 of the body where it senses blood temperature affectedby the energy source 210. The sensor 200 can be located, for example, inthe arm 104, leg or neck. The sensor 200 and the sensor 20 shown in FIG.1 can be similar in construction and are similar in function.

[0059] When the system is configured as shown in FIG. 11, cardiac outputand blood flow may be calculated as previously described. When thesensor 200 and the energy source 210 are separated by such a distancethat a material amount of thermal energy from the energy source is lostto the body before reaching sensor 200, a compensating factor isincluded in the calibration constant referenced in FIG. 5.

[0060] In some situations it may be desirable to provide more frequentindications of cardiac output, such as, for example, the instantaneouscardiac output or the cardiac output averaged over each individualcardiac cycle (i.e. each heart beat). Such information can be providedusing the further embodiments of the invention discussed below withreference to FIGS. 7-8. A single temperature sensor 30 at a locationnear the distal end of the catheter 31 (as shown in FIG. 7A) can be usedto determine the instantaneous or beat-to-beat blood velocity V(t). Theblood velocity can be combined with the cardiac output averaged over oneor more data collection cycles to calculate instantaneous cardiacoutput. The process used is shown in the process depicted in FIG. 7B,identified as Process IV.

[0061] As seen therein, the initial temperature T_(3i)(t) sensed attemperature sensor 30 as a function of time is smoothed, or filtered, inthe manner as previously discussed above, and suitably measured andstored at an initial time t₀. A predetermined rise in temperature ΔT₃ ofthe temperature sensor itself is selected. Power is then supplied attime t₀ to the temperature sensor 30 from a power source 30A connectedthereto to cause its temperature T₃(t) to rise by a predeterminedamount.

[0062] Power may be supplied to the sensor in different ways accordingto the needs of the particular measurement and the relative simplicityor complexity of the required circuitry, three such ways being depictedin FIGS. 8A, 8B, and 8C.

[0063] For example, in a first mode of operation (FIG. 8A), heatingpower may be supplied to the sensor in such a manner as to keep thefinal sensor temperature T_(3f)(t) constant at an initial level ΔT₃above the initial temperature T_(3i)(t₀), i.e. T_(3f)=T_(3i)(t₀)+ΔT₃even when the local blood temperature varies with time, as illustratedin FIG. 9A. Under such conditions, the sensor is maintained at atime-varying temperature increment ΔT₃(t) above the instantaneous localblood temperature, T_(b)(t).

[0064] Alternatively, in a second mode of operation, power can besupplied to the sensor so as to continuously maintain the sensor at afixed temperature increment above the time varying local bloodtemperature, as illustrated in FIG. 9B. Under such conditions,ΔT₃(t)=ΔT₃, a constant, and the sensor temperature varies according toT_(3f)(t)=ΔT₃+ΔT_(3i)(t).

[0065] A third mode of heating may also be convenient when thetemperature sensors are temperature-sensitive resistors, or thermistors.Thus, when a thermistor is used, it may be more convenient to design anelectrical heating circuit that maintains the sensor at a constantresistance increment above the resistance of the sensor that correspondsto the local blood temperature. If R is the corresponding resistance fora sensor temperature T, then these conditions are represented byΔR₃(t)=ΔR₃, a constant, and the sensor resistance varies according toR_(3f)(t)=ΔR₃+ΔR_(3i)(t), as illustrated in FIG. 9C. The change intemperature ΔT₃(t) is then replaced by the change in resistance R₃(t) inthe ratio which is integrated over a cardiac cycle. Further details andexemplary apparatus for such modes of operation are presented anddescribed in U.S. Pat. No. 4,059,982, issued to E. F. Bowman on Nov. 29,1977. With all three of the above approaches, power (P) is supplied toproduce a temperature rise (ΔT) both of which are then related to theinstantaneous blood velocity and, hence, blood flow.

[0066] Techniques in which sensor heating power and temperature can bemeasured and used to provide more detailed information on cardiac outputare described below. The technique involved can be applied to measureboth instantaneous cardiac output, and the cardiac output for anindividual cardiac cycle. Such detailed measurement information greatlyenhances the diagnostic capability of a physician.

[0067] First, a method is described to measure instantaneous volumetricflow (which flow if measured at the location described above is thecardiac output). For each of the particular implementations describedabove, the power P₃(t) applied to the temperature sensor 30 iscontrolled so as to maintain the final temperature of the sensor at adesired value T_(3f). The power applied to the temperature sensor 30 or,more generally, the ratio of the power applied to the sensor to thetemperature increment, P₃(t)/ΔT₃(t), is directly correlated with thefluid and flow properties of the flowing liquid about the sensor.

[0068] For example, the relationship between required sensor power andlocal fluid velocity, V(t), is given by a correlation of the form:

P(t)=4πkaΔT(t)[1+C ₁ P _(r) ^(n)(2aρV(t)/μ)^(m)]

[0069] Where

[0070] P(t)=instantaneous power to sensor

[0071] k=thermal conductivity of fluid

[0072] a=sensor radius

[0073] ΔT(t)=instantaneous temperature difference between heated sensorand unheated fluid temperature.

[0074] C₁=constant of calibration

[0075] P_(r)=a non-dimensional “Prandtl” number which relates to theviscosity μ, heat capacity Cp and thermal conductivity k of a fluid.

[0076] n,m=power factors which are determined from experimental data

[0077] ρ=fluid density

[0078] μ=viscosity

[0079] V(t)=instantaneous fluid velocity

[0080] The fluid flow velocity in the vicinity of the sensor can bedetermined from the required sensor heating power. Volumetric flow inthe vessel can then be determined with one further assumption for thedistribution of the fluid flow within the vessel. For example, assuminga uniform velocity profile within the vessel, volumetric flow F₃ isgiven by

F₃=V A

[0081] where V is the fluid velocity in the vessel and A is the flowarea.

[0082] If the fluid flow area A is not previously known, it may beinferred from the measurement of average volumetric flow in the vessel.Such average volumetric flow can be determined, for example, by usingthe techniques of the invention already described above herein or byusing other techniques for yielding comparable information. For example,if F is the average cardiac output, typically measured over severalcardiac cycles, as described above, and V is the average fluid velocity,determined by calculating an average value for the instantaneous flowvelocity over at least one cardiac cycle, then one such estimate for theaverage flow area A is given by

{overscore (A)}={overscore (F)}|{overscore (V)}

[0083] Therefore, given the sensor measured heating power, first thefluid velocity and then volumetric flow can be calculated at any desiredinstant in time, i.e., F_(3(t))=V(E){overscore (A)}, yielding aninstantaneous measure of volumetric flow, i.e., cardiac output.

[0084] In another embodiment, a method to measure cardiac output over asingle cardiac cycle is described. As described above in differentimplementations, the power P₃(t) applied to the temperature sensor 30 iscontrolled so as to maintain the temperature of the sensor at a desiredsignal value T_(3f). The power applied to the temperature sensor 30, ormore generally, as discussed above, the ratio of the power applied tothe sensor to the temperature increment, i.e., P₃(t)/ΔT₃(t), is directlycorrelated with the properties of the fluid flow in the vicinity of thesensor.

[0085] Thus, the integrated value of the power to temperature ratio overa single cardiac cycle is directly correlated with, i.e., isproportional to the average cardiac output over the cardiac cycle,${\int_{{cardiac} - {cycle}}{\frac{P_{3}(t)}{\Delta \quad {T_{3}(t)}}{t}}} \propto \overset{\_}{F}$

[0086] or, alternatively expressed${\frac{1}{T}{\int_{{cardiac} - {cycle}}{\frac{P_{3}(t)}{\Delta \quad {T_{3}(t)}}{t}}}} \propto \overset{\_}{F}$

[0087] where T represents the period of the cardiac cycle and Fcorrespondingly represents cardiac output averaged over the cardiaccycle. Thus, the average cardiac output F over an individual cardiaccycle then can be determined from the measured and integrated power andtemperature signals from the sensor.

[0088] Furthermore, an explicit correlation for integrated power andaverage cardiac output over the cardiac cycle may be dispensed with if asimple qualitative indication of the change in cardiac output on acardiac cycle-to-cardiac cycle basis is desired. To obtain suchinformation, a given measurement of cardiac output is taken asassociated with a corresponding measured value of the integrated sensorsignal over a cardiac cycle. The measurement of cardiac output could beobtained intermittently by the techniques described in this invention orother similar techniques. Since cardiac output is known to be correlatedwith the value of the integrated sensor signal over the cardiac cycle,any changes in the sensor signal indicate a corresponding change incardiac output.

[0089] In certain situations, it may be desirable to compensate fortemperature variations in the blood which is flowing past the sensor, asthis may affect the value of F(t). A process for such compensation isdepicted as Process IV in FIG. 7B wherein a temperature T₂(t) is sensedby a second sensor (which may be, for example, sensor 19 or sensor 20)at a location remote from sensor 30 (see FIG. 7A). For example,knowledge of the instantaneous blood temperature is required for theprocess in which the heated sensor is maintained at a constant incrementabove the local blood temperature. In this case, the temperature T₂(t)is used as a proxy for the temperature T₃(t) which would be measured inthe absence of sensor heating.

[0090] In a further alternative embodiment, where only two sensors 19and 20 are utilized (as shown in FIG. 2), sensor 20 can be used as theprimary sensor when calculating instantaneous cardiac output (equivalentto sensor 30 in FIG. 7A) and sensor 19 can be used as the secondarytemperature compensation sensor. In such an embodiment, the averagedcardiac output can be determined using sensors 19 and 20, as set forthin FIGS. 3-6 and the instantaneous cardiac output can then subsequentlybe determined using sensors 19 and 20, as set forth in FIG. 7B andeither FIGS. 8A, 8B or 8C, such average and instantaneous cardiac outputdeterminations being made in sequence by the microprocessor to providethe cardiac information in both forms, as desired.

[0091] As mentioned above, when using the above described catheter, thevarious flow values which are determined in accordance with theprocesses as discussed above are proportional to flow but may not beequal to the actual flow values unless they are suitably calibratedsince the correspondence between the calculated and actual valuesdepends on the manner in which a particular catheter is constructed andused. A calibration constant for a particular catheter can berepresented by the slope and intercept of a curve which relates thecalculated flow and the actual flow, in accordance with the followingrelationship:

F _(actual) =aF _(calc) +b

[0092] where, as illustrated in FIG. 10, “a” is the slope of a straightline 35 and “b” is the intercept thereof along the vertical axis. Curve35 can be obtained by using a known catheter and known flow valuestherein to construct a curve 36. The best straight line fit isdetermined as line 35. The slope “a” and intercept “b” are therebydetermined. Such determined values for “a” and “b” can be used with thecalculated flow values in each case to determine the actual flow fromthe calculated flow. While the above description discusses preferredembodiments of the invention, modifications thereof may occur to thosein the art within the spirit and scope of the invention. Hence, theinvention is not to be construed as limited to particular embodimentsdescribed, except as defined by the appended claims.

1. A system for quantifying blood flow in a living subject comprising: (a) a thermal energy source for increasing blood temperature at a location in the venous side of the circulatory system during preselected time intervals; (b) a first sensor for sensing blood temperature in the venous side of the circulatory system where blood temperature is substantially unaffected by the output of the thermal energy source; (c) a second sensor for sensing blood temperature in the arterial blood flow path where the blood temperature is affected by the output of the thermal energy source; and (d) means responsive to the first and second sensors and output of said thermal energy source for calculating a blood flow related value as a function of outputs of the first and second sensors when blood temperature in the arterial side of the circulatory system is affected by output of said thermal energy source and the outputs of the first and second sensors when blood temperature in the arterial side of the circulatory system is not substantially affected by said thermal energy source.
 2. A system according to claim 1 further comprising a catheter adapted to be introduced into the venous system for supporting one or both of the first sensor and the thermal energy source.
 3. A system according to claim 2 further comprising a catheter adapted to be introduced into the arterial system for supporting the second sensor.
 4. A system according to claim 3 wherein said calculating means comprises means for (i) determining the difference in the temperatures sensed by the first and second sensors when the blood temperature in the arterial side of the circulatory system is affected by output of the thermal energy source; (ii) determining the difference in the temperatures sensed by the first and second sensors when the blood temperature in the arterial side of the circulatory system is substantially unaffected by output of the thermal energy source and (iii) calculating blood flow as a function of the difference between said differences.
 5. A system according to claim 1 wherein said calculating means comprises means for calculating blood flow as a function of the output of the thermal energy source and the difference between the temperature difference sensed by the first and second sensors when the blood temperature in the arterial side of the circulatory system is affected by output of the thermal energy source and the temperature difference sensed by the first and second sensors when the thermal energy source is not activated.
 6. A system according to claim 1 wherein the thermal energy source comprises an energy source for increasing blood temperature in or near the right atrium or vena cava and said calculating means calculates a blood flow value that corresponds to cardiac output.
 7. A system for quantifying blood flow in the circulatory system of a living subject comprising: (a) heating means adapted to be located in the venous flow path for intermittently elevating the temperature of blood in the arterial flow path; (b) means for sensing blood temperature in the arterial flow path; (c) means for providing a value corresponding to blood temperature at a location in the venous flow path not affected by the output of said heating means; and (d) means for calculating blood flow as a function of the output of said heating means when elevating blood temperature, the difference in the values from said sensing means and said providing means when the temperature of blood is not elevated by said heating means and the difference in the values from said sensing means and said providing means when the temperature of blood is elevated by said heating means.
 8. A system according to claim 7 wherein said providing means comprises a second means for sensing blood temperature.
 9. A method for calculating blood flow in the blood flow path of a living subject comprising: (a) changing a thermal energy characteristic of blood at a site in the venous side of the circulatory system in or near the right atrium or vena cava during preselected time intervals; (b) detecting the temperature difference between a location in the venous system substantially unaffected by said changes introduced at the site and a selected location in the arterial system where blood temperature is affected by said changing step; (c) detecting the temperature difference between the location in the venous system and the location in the arterial system in the absence of said changing step; and (d) calculating blood flow as a function of the temperature difference of the first said detecting step, the temperature difference of the second said detecting step and the change introduced by said changing step. 